The goal of drug delivery systems is to increase the efficacy and safety of both new and existing drugs. A number of drug compounds cannot be delivered safely and/or effectively by conventional routes or dosage forms such as oral tablets or injection. Alternative delivery methods can increase safety by sequestering reagents in carriers that reduce systemic exposure and decrease dose-limiting toxicity and side effects, or by providing sustained delivery so that therapeutic levels can be achieved with fewer and smaller doses. New delivery systems can also increase efficacy by several strategies, including:
Increasing stability of the drug;
Increasing the ability of the drug to reach its therapeutic target by prolonging the circulating half-life; and
Targeting delivery to the therapeutic site in order to reduce the total circulating dose without diminishing efficacy;
For drugs with a narrow therapeutic index, many of the approaches that reduce toxicity also enhance efficacy thereby increasing the therapeutic window and efficacy of drugs. The most prominent delivery systems used in the clinic are based on liposomes (e.g. Doxil which are liposomes filled with doxorubicin).
There are two liposomal targeting strategies: passive and specific. Passive targeting refers to the preferential accumulation of liposomes in tumours and at sites of infection and inflammation. Small sterically stabilized liposomes extravasate through leaky blood vessels that are formed through tumour angiogenesis or damaged by infection and inflammation. The liposomes accumulate in tumour interstices and at sites of infection and inflammation, where they gradually release their encapsulated drugs. This sequesters potentially toxic agents from susceptible non-target sites such as the brain, liver and heart. The mechanism of liposome accumulation may be a combination of the leakiness of the newly forming or damaged capillaries and enhanced vascular permeation by the coated liposomal particles themselves.
Specific targeting involves the use of antibodies or ligands to tag liposomes so that they bind specifically to cells that express the appropriate cell-surface antigens or ligand receptors, respectively. In principle, liposomes can be targeted to any cell surface structure that can be recognised by a fragment of a specific antibody, or to any receptor for which a small and specific ligand can be produced. Hence, liposomes can be directed to specific classes of T and B lymphocytes or to tumour cells preferentially expressing high levels of specific cell surface proteins. The goals of ligand targeting of liposomes are to concentrate them selectively at the therapeutic site, decrease the required dose by reducing non-specific losses, and reduce systemic exposure to reagents with toxic side effects. There are two caveats, however: 1) The target cells must be accessible; and 2) The reagents must be released from the liposomes efficiently enough to have a clinically significant effect.
One important caveat of liposomal drug delivery is the efficiency of drug release from the liposome at the target location, which determines the clinical effect. This release can be governed by multiple processes and variables. Localisation of passive or active targeted liposomes is usually followed by a relatively lengthy process which can involve an internalisation pathway followed by intracellular processing and drug release or expulsion back to the extracellular domain. Alternatively, these constructs stay in the extracellular domain and are slowly degraded, depending on the specific environment. When liposomes are taken up by cells through endocytosis, the liposome needs to be degraded and the drug has to be able to escape the lysosomal compartment. In this respect it should be noted that endothelial cells do not possess a machinery to degrade liposomes and as a result these constructs are expulsed back to the extracellular environment. Currently the targeting and localisation of liposomes can be controlled by passive or active methods, but there is only an indirect, unpredictable and non-universal control over the subsequent release from these liposomes. The release pathway, the extent of release, the release profile (slow or bolus), and the timing, all depend on the specifics and peculiarities of the target.
To address this issue, liposomal delivery systems capable of near instantaneous release of their content under the influence of redox state, pH or light have been developed.
The technology is based on a well-studied bacterial channel protein “Mechanosensitive channel of large conductance”, MscL, from E. coli. In its native form the channel creates a large non-selective pore of 3-4 nm in diameter in the membrane and allows the passage of not only ions but also small molecules, peptides and smaller proteins (up to 7 kDa). In nature, MscL opens in response to the tension in the membrane. It has been shown that the hydrophilicity of the 22nd amino acid position of MscL affects the mechanosensitivity of the channel up to a point where it starts to open even in the absence of tension (Yoshimura et al. (1999) Biophys. J. 77, 1960-1972). Hydrophilic substitutions in this narrow pore constriction area of the channel cause hydration of the pore and weakening of the hydrophobic van der Waals forces responsible for the close packing of the inner membrane helices in the closed state of the channel. The effect is reinforced if charged or bulky groups are introduced because of electrostatic repulsion and steric factors, respectively. This is reflected in the energetics of the gating transitions and leads to the opening of the channel even in the absence of tension.
On the basis of this principle, the MscL protein was re-engineered to site-selectively-incorporate (masked) amine-functionalised molecules. A series of small modulators were designed, synthesised and specifically attached to an engineered Cys at position 22 in MscL. The working principle, depicted in FIG. 1, is that the protein-attached modulators would be charged only in response to a pre-defined stimulation (pH, light, etc) leading to hydration of the hydrophobic constriction zone of the pore and channel opening in the absence of the natural stimulus. The masked reagents possess a nitrophenol moiety, which is removed upon illumination. This affords a free amino moiety which, depending on the pH, can get protonated and trigger the opening of the channel. The ability to control the release of liposome content with reversible channel opening and closing was demonstrated under the influence of UV and visible light, respectively (Kocer et al. (2005) Science 309, 755-758) and in response to a decrease in pH using channels modified to respond directly to pH as well as channels engineered (using masked reagents) to respond to pH only after illumination (Kocer et al. (2006) Angew. Chem. Int. Ed. 45, 3126-3130). Rationally designed chemical modulators convert a bacterial channel protein into a pH-sensory valve. This methodology is also disclosed in PCT patent applications WO2005051902, WO03084508 and WO03000233.
The activateable liposomal drug delivery systems discussed above allow an increased level of control over drug release from liposomes using light- or pH-mediated release. These mechanisms provide an additional selectivity for a specific local environment (low pH in certain tumours) or localised illumination. However, the drawback of this additional selectivity is that these tools can not be universally applied. Due to the low penetration depth of light, the technology of controlled drug delivery is limited to disorders situated at or near the body surface or in combination with a catheter-light pipe. With respect to the pH-activateable liposomes, many potential targets do not have a pH that is significantly different from surrounding non-target tissue.
Accordingly, there remains a need for drug delivery systems whereby the drug release can be effectively controlled.